The present invention relates generally to diagnostic imaging and, more particularly, to a direct conversion detector capable of providing photon count and/or energy data with reduced charge sharing between pixels of the direct conversion detector.
Typically, in radiographic imaging systems, such as x-ray and computed tomography (CT), an x-ray source emits x-rays toward a subject or object, such as a patient or a piece of luggage. Hereinafter, the terms “subject” and “object” may be interchangeably used to describe anything capable of being imaged. The beam, after being attenuated by the subject, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the x-rays. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
In some CT imaging systems, for example, the x-ray source and the detector array are rotated within a gantry and within an imaging plane around the subject. X-ray sources for such CT imaging systems typically include, but are not limited to, x-ray tubes, solid state x-ray source, thermionic x-ray sources, and field emitters which emit the x-rays as a fan beam emanating from a focal point. X-ray detectors for such CT imaging systems are typically configured in a circular arc centered to a focal spot. In addition, such detectors include a collimator for collimating x-ray beams received at the detector which focus to a focal spot. Such detectors include a scintillator for converting x-rays to light energy adjacent the collimator and a photodiode for receiving the light energy from an adjacent scintillator and producing electrical signals therefrom. Typically, each scintillator of a scintillator array converts x-rays to light energy. Each photodiode detects the light energy and generates a corresponding electrical signal as a function of the light emitted by a corresponding photodiode. The outputs of the photodiodes are then transmitted to the data processing system for image reconstruction.
Conventional CT imaging systems utilize detectors that convert radiographic energy into current signals that are integrated over a time period, then measured and ultimately digitized. A drawback of such detectors however is their inability to provide data or feedback as to the number and/or energy of photons detected. That is, conventional CT detectors have a scintillator component and photodiode component wherein the scintillator component illuminates upon reception of radiographic energy and the photodiode detects illumination of the scintillator component and provides an electrical signal as a function of the intensity of illumination. Under the charge integration operation mode, the photodiode is not capable of discriminating between the energy level or the photon count from the scintillation. For example, two scintillators may illuminate with equivalent intensity and, as such, provide equivalent output to their respective photodiodes. Yet, the number of x-rays received by each scintillator may be different as well as the x-rays intensity, but yield an equivalent light output.
Energy discriminating, direct conversion detectors are capable of not only x-ray counting, but also providing a measurement of the energy level of each x-ray detected. Consequently, such a detector could potentially be used for SPECT or PET imaging. Energy discriminating detectors can also be used to give compositional information of an imaged object by applying a material discrimination algorithm on measured energy levels. While a number of materials may be used in the construction of a direct conversion energy discriminating detector, semiconductors such as typically Cadmium Zinc Telluride (CZT), Cadmium Telluride (CdTe) and the like have been shown to be preferred materials.
In a typical imaging application, x-rays are absorbed in the direct conversion material which results in creation of an electrical charge in the direct conversion material. In order to create digital image information, the charge generated is collected on segmented anodes typically using either charge integration or charge pulse counting electronics.
A drawback of direct conversion semiconductor detectors, however, is that x-rays absorbed in the direct conversion material near the gaps or perimeters of the anodes can result in a charge being generated therein that is shared by at least two neighboring pixel anodes. When using charge integration electronics, charge sharing can manifest itself as crosstalk between neighboring pixels, thus rendering the electronics susceptible to electronic noise amplification and spatial blurring of the image. When using pulse counting electronics, charge sharing can result in dividing the charge between at least two anodes, resulting in lost counts when the amplitude of the charge pulse collected in at least one of the anodes is below a discrimination threshold. Additionally, when pulse counting, high energy x-rays can result in loss of detection quantum efficiency (DQE) by the creation of two or more counts being collected in two or more neighboring anodes, thus mis-counting the events and binning, for instance, a single high energy event as two or more low-energy events. The mis-counting of events and binning with respect to energy will degrade the capability for material discrimination.
Another drawback of direct conversion semiconductor detectors with regard to CT imaging is that the response at the edge and corners of the direct conversion crystal is not reproducible. Such locations of a direct conversion crystal typically have charge trapping centers that cause changes in the internal electric field as the incident x-ray flux changes. The changing internal field can cause a poor detector response that can lead to a non-optimal image.
Another drawback of direct conversion semiconductor detectors with regard to CT imaging is that these types of detectors cannot count at the very high x-ray photon flux rates typically encountered with conventional CT systems, e.g., at or above 5-100 million counts per sec per millimeter squared (Mcps). The very high x-ray photon flux rate causes pile-up and polarization, which ultimately can lead to detector saturation. That is, these detectors typically saturate at relatively low x-ray flux level thresholds. Above these thresholds, the detector response is not predictable or has degraded dose utilization. Saturation can occur at detector locations wherein small subject thickness is interposed between the detector and the radiographic energy source or x-ray tube. It has been shown that these saturated regions correspond to paths of low subject thickness near or outside the width of the subject projected onto the detector fan-arc. In many instances, the subject is more or less circular or elliptical in the effect on attenuation of the x-ray flux and subsequent incident intensity to the detector. In this case, the saturated regions represent two disjointed regions at extremes of the fan-arc. In other instances, saturation occurs at other locations and in more than two disjointed regions of the detector. In the case of an elliptical subject, the saturation at the edges of the fan-arc is reduced by the imposition of a bowtie filter between the subject and the x-ray source. Typically, the filter is constructed to match the shape of the subject in such a way as to equalize total attenuation, filter and subject, across the fan-arc. The flux incident to the detector is then more closely uniform across the fan-arc and does not result in saturation. However, the bowtie filter may not be optimal given that a subject population is significantly less than uniform and not exactly elliptical in shape. In such cases, it is possible for one or more disjointed regions of saturation to occur or, conversely, to over-filter the x-ray flux and create regions of very low flux. Low x-ray flux in the image projection tends to increase noise in the reconstructed image of the subject.
Detector saturation causes loss of imaging information and results in artifacts in x-ray projection and CT images. In addition, hysteresis and other non-linear effects occur at flux levels near detector saturation as well as at flux levels over detector saturation. Direct conversion detectors are susceptible to a phenomenon called “polarization,” where charge trapping inside the material changes the internal electric field, alters the detector count and energy response in an unpredictable way, and results in hysteresis where response is altered by previous exposure history. In particular, photon counting, direct conversion detectors saturate due to the intrinsic charge collection time (i.e., dead time) associated with each x-ray photon event. Saturation will occur due to pulse pile-up when the x-ray photon absorption rate for each pixel is on the order of the inverse of the charge collection time.
A number of techniques have been developed to address charge-sharing in direct conversion detectors. Energy discriminating detectors typically comprise a number of segmented anodes that define a pixellated structure onto which the direct conversion material is electrically attached. The anodes define the response area of the imaging pixels which segment the area of the detection plane. Smaller pixels are generally desirable because they make available higher spatial resolution information which can result in higher resolution images and because the flux rate capability is generally improved with smaller pixels. However, smaller pixel size can result in higher cost because there are more channels per unit area which need to be connected to readout electronics. In addition, smaller pixels or detector elements have larger perimeter-to-area ratios resulting in a larger percentage of charge sharing regions per unit area of the detector.
Because the perimeters of the pixels is the region where a charge may be shared between two or more pixels, incomplete energy information and/or a miscount of x-ray photons occurs for such a charge because the readout electronics are not configured to combine near-simultaneous signals in neighboring pixels. Readout electronics could incorporate a time-coincidence circuit configured to identify events occurring within a defined time window that, once identified, prevents the detected event from receiving a bin count. However, such electronics can be costly and difficult to implement. A time-coincidence circuit would also not adequately preserve energy information about the x-ray event shared between two or more pixels without suffering degradation due to chance coincidence occurring with the near simultaneous arrival of two or more photons in neighboring regions.
To solve the problem regarding the reproducibility of the response at the edges and corners of the direct conversion crystal, a guard ring is typically placed on the anode surface of the device or on side walls of the crystal walls. However, a guard ring does not prevent trapping of charge within the semiconductor, and a guard ring does not prevent a changing electric field from developing within the semiconductor.
Therefore, it would be desirable to design a CT apparatus and method to reduce charge sharing between pixels of the direct conversion detector.